Post electron beam conditioning of polymeric medical devices

ABSTRACT

Methods are disclosed for conditioning a polymeric stent after sterilization, and/or after crimping and before packaging, such that the properties of the polymeric stent fall within a narrower range of values. The stent is exposed to a controlled temperature at or above ambient for a period of time after radiation sterilization and/or after crimping and before sterilization. As a result, the polymeric stent properties, particularly radial strength and number-average molecular weight of the polymer of the polymeric stent, fall within a narrower range.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.14/139,740 filed on Dec. 23, 2013 which is a divisional of U.S.application Ser. No. 13/093,755 filed on Apr. 25, 2011 which is acontinuation-in-part of co-pending U.S. application Ser. No. 12/860,681,filed on 20 Aug. 2010, and is also a continuation-in-part of co-pendingU.S. application Ser. No. 12/764,803, filed on 21 Apr. 2010. All ofthese applications are incorporated by reference in their entirety,including any drawings herein.

BACKGROUND OF THE INVENTION

Field of the Invention

This invention relates to methods making stents from bioabsorbablepolymers.

Description of the State of the Art

This invention relates to radially expandable endoprostheses that areadapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel. Astent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices that function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of a bodily passage or orifice. In suchtreatments, stents reinforce body vessels and prevent restenosisfollowing angioplasty in the vascular system. “Restenosis” refers to thereoccurrence of stenosis in a blood vessel or heart valve after it hasbeen treated (as by balloon angioplasty, stenting, or valvuloplasty)with apparent success.

Stents are typically composed of scaffolding that includes a pattern ornetwork of interconnecting structural elements or struts, that may beformed from wires, tubes, or sheets of material rolled into acylindrical shape. This scaffolding gets its name because it physicallyholds open and, if desired, expands the wall of the passageway, orlumen. Typically, stents are capable of being compressed, or crimped,onto a catheter so that they can be delivered to and deployed at atreatment site. Delivery includes inserting the stent through smalllumens, such as blood vessels, using a catheter and transporting it tothe treatment site. Deployment includes expanding the stent to a largerdiameter once it is at the desired location. Mechanical interventionwith stents has reduced the rate of restenosis as compared to balloonangioplasty. Yet, restenosis remains a significant problem. Whenrestenosis does occur in the stented segment, the treatment of it can bechallenging, as clinical options are more limited than for those lesionsthat were treated solely with a balloon.

Stents are used not only for mechanical intervention but also asvehicles for providing biological therapy. Biological therapy usesmedicated stents to locally administer a drug. A medicated stent may befabricated by coating the surface of either a metallic or polymericscaffolding with a polymeric carrier that includes a drug. Polymericscaffolding may also serve as a carrier of a drug.

It may be desirable for a stent to be biodegradable. In many treatmentapplications, the presence of a stent in a body may be necessary for alimited period of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished.Therefore, stents fabricated from biodegradable, bioabsorbable, and/orbioerodable materials such as bioabsorbable polymers should beconfigured to completely erode only after the clinical need for them hasended.

One of the challenges of making medical devices out of polymers is thatthe properties of a polymer can change both during processing and afterprocessing. These properties include mechanical properties such asstrength and toughness as well as bioresorption kinetics. The processingsteps in a fabrication process of a stent may be designed to maintain orinstill in the stent particular ranges of the strength, toughness, andbioresorption, that are crucial for treatment with the stent. In somecases, properties of the polymer can change during additional processingoperations and/or as a function of time during storage. Therefore,methods are needed that reduce, or eliminate undesirable changes inproperties, and/or ameliorate their impact.

SUMMARY OF THE INVENTION

Various embodiments of the present invention include a method forconditioning a polymeric stent. The method may include the operationsof: exposing a polymeric stent with a polymeric scaffolding to atemperature equal to, approximately equal to, or greater than 30° C. andnot more than about 15° C. less than the glass transition temperature ofthe polymeric scaffolding for a duration of time. The duration of timemay be at least 8 hours. The polymeric stent may have been crimped ontoa delivery device, packaged, and sterilized prior to the exposure. Thepolymeric scaffolding may be formed from a polymeric article that hasbeen deformed by the application of stress at a temperature greater thanthat of the glass transition temperature of the polymeric article, andthe polymeric article may have a glass transition temperature greaterthan 25° C. The exposure temperature may be controlled to within ±3° C.

In an aspect of the invention, the polymeric article may be a polymertube and the deformation under stress may comprise radial expansion ofthe polymer tube.

In another aspect of the invention, the polymeric scaffolding mayinclude a polymer selected from the group consisting of poly(L-lactide),polymandelide, poly(DL-lactide), polyglycolide,poly(L-lactide-co-glycolide), and all combinations thereof in allproportions.

In another aspect of the invention, the exposure temperature may be nothigher than 20° C. below the glass transition temperature of thepolymeric scaffolding.

In another aspect of the invention, the duration of exposure may be fromabout 8 hours to about 20 days and the exposure temperature may be fromabout 32° C. to about 40° C.

In other aspects of the invention, the duration of exposure may be fromabout 1 day to about 10 days, or from about 2 days to about 6 days.

In another aspect of the invention, the exposure temperature may be inthe range of about 35° C. to about 40° C.

In another aspect of the invention, the method may also include exposingthe polymeric stent to a temperature equal to or greater than 35° C. andnot more than about 10° C. greater than the glass transition temperatureof the polymeric scaffolding for a duration of time where the durationof time may be in the range of from about 4 hours to about 10 days, andwhere the exposure may occur after the polymeric stent has been crimpedonto a delivery device, but before the polymeric stent has beensterilized. The temperature of the exposure after crimping and beforesterilization may be controlled to within ±3° C.

In another aspect of the invention, the duration of the exposure aftercrimping and before sterilizing may be from about 16 hours to about 48hours, and the temperature of the exposure after crimping and beforesterilization is from about 45° C. to about 65° C.

In another aspect of the invention, the duration of the exposure aftercrimping and before sterilizing may be from about 16 hours to about 32hours, and the temperature of the exposure after crimping and beforesterilization may be from about 50° C. to about 65° C.

In another aspect of the invention, the polymeric stent may be crimpedonto a delivery device and the crimping may be performed at atemperature in the range of about 45° C. to about 50° C.

In another aspect of the invention, the polymeric stent may be crimpedin the range of about 40° C. to about 55° C., and the post-sterilizationexposure temperature may be about 33° C. and not more than about 37° C.and the duration of the exposure may be in the range of about 32 hoursto about 84 hours.

In another aspect of the invention, the polymeric stent may be crimpedonto a delivery device and the crimping may be performed at atemperature in the range of about 48° C., and the post-sterilizationexposure temperature may be about 35° C. and the duration of theexposure may be in the range of about 48 hours to about 72 hours.

Various embodiments of the present invention include a method forconditioning a polymeric stent. The method includes the operations of:exposing a polymeric stent with a polymeric scaffolding consistingessentially of poly(L-lactide) to a temperature equal to, approximatelyequal to, or greater than 30° C. and not more than about 55° C. for aduration of time. The duration of time may be at least 8 hours. Thepolymeric stent may have been crimped onto a delivery device, packaged,and sterilized prior to the exposure. The polymeric scaffolding may beformed from a polymeric tube consisting essentially of poly(L-lactide)that has been deformed by the radial expansion of the polymeric tube ata temperature greater than that of the glass transition temperature ofthe polymeric tube. The exposure temperature may be controlled to within±3° C.

In an aspect of the present invention, the duration of exposure aftersterilization may be from about 8 hours to about 20 days and theexposure temperature may be from about 32° C. to about 40° C.

In an aspect of the present invention, the duration of exposure aftersterilization may be from about 1 day to about 10 days.

In an aspect of the present invention, the duration of exposure aftersterilization may be from about 2 days to about 6 days, and wherein theexposure temperature may be in the range of about 35° C. to about 40° C.

An aspect of the present invention, the method of exposing apoly(L-lactide) scaffolding also may include exposing the polymericstent with the polymeric scaffolding to a temperature equal to orgreater than 35° C. and not more than about 70° C. for a duration oftime, where the duration of time may be from about 4 hours to about 6days, after the polymeric stent has been crimped onto a delivery device,but before the polymeric stent has been sterilized. The temperature ofthe exposure after crimping and before sterilization may be controlledto within ±3° C.

In an aspect of the invention, for a polymeric stent with apoly(L-lactide) scaffolding, the duration of the exposure after crimpingand before sterilizing may be from about 16 hours to about 48 hours, andthe temperature of the exposure after crimping and before sterilizationmay be from about 45° C. to about 65° C.

In an aspect of the invention, for a polymeric stent with apoly(L-lactide) scaffolding, the duration of the exposure after crimpingand before sterilizing may be from about 16 hours to about 32 hours, andthe temperature of the exposure after crimping and before sterilizationmay be from about 50° C. to about 65° C.

In an aspect of the invention, for a polymeric stent with apoly(L-lactide) scaffolding, the polymeric stent may be crimped onto adelivery device at a temperature in the range of about 45° C. to about50° C.

In an aspect of the invention, for a polymeric stent with apoly(L-lactide) scaffolding, the post-sterilization exposure temperaturemay be in the range from about 33° C. and not more than about 37° C.,and the duration of the exposure may be in the range of about 32 hoursto about 84 hours.

In an aspect of the invention, for a polymeric stent with apoly(L-lactide) scaffolding, the polymeric stent may be crimped onto adelivery device at a temperature of about 48° C., and thepost-sterilization exposure temperature may be about 35° C., and theduration of the exposure after sterilization may be in the range ofabout 48 hours to about 72 hours.

Various embodiments of the present invention include a method forconditioning a polymeric stent. The method may include the operationsof: exposing a polymeric stent with a polymeric scaffolding to atemperature equal to, approximately equal to, or greater than 30° C. andnot more than about 55° C. for a duration of time until the radialstrength is reduced by at least 10%. The duration of time may be atleast 30 minutes, and the polymeric stent may have been crimped onto adelivery device, packaged, and sterilized prior to the exposure. Thepolymeric scaffolding may be formed from a polymeric tube that has beendeformed by the radial expansion of the polymeric tube at a temperaturegreater than that of the glass transition temperature of the polymerictube, and the exposure temperature may be controlled to within ±3° C.The glass transition temperature of the polymeric scaffolding may begreater than 25° C.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts an exemplary stent.

FIG. 2 depicts the specific volume of a polymer as a function oftemperature.

FIG. 3 depicts the relative free radical concentration of a polymericstent with heat treatment and the stent with no heat treatment.

FIG. 4 depicts the relative free radical concentration of a polymericstent with heat treatment and the stent with no heat treatment.

FIG. 5 depicts the number-average molecular weight of a polymer of apolymeric stent as a function of time after electron beam sterilization.

FIG. 6 depicts the radial strength of a polymeric stent as a function oftime after electron beam sterilization.

FIG. 7 depicts the heat capacity of a polymer as a function oftemperature.

DETAILED DESCRIPTION OF THE INVENTION

Use of the term “herein” encompasses the specification, the abstract,and the claims of the present application.

Use of the singular herein includes the plural and vice versa unlessexpressly stated to be otherwise, or obvious from the context that suchis not intended. That is, “a” and “the” refer to one or more of whateverthe word modifies. For example, “a drug” includes one drug, two drugs,etc. Likewise, “the polymer” may refer to one, two or more polymers, and“the device” may mean one device or a plurality of devices. By the sametoken, words such as, without limitation, “polymers” and “devices” wouldrefer to one polymer or device as well as to a plurality of polymers ordevices unless, again, it is expressly stated or obvious from thecontext that such is not intended.

As used herein, unless specifically defined otherwise, any words ofapproximation such as without limitation, “about,” “essentially,”“substantially,” and the like mean that the element so modified need notbe exactly what is described but can vary from the description. Theextent to which the description may vary will depend on how great achange can be instituted and have one of ordinary skill in the artrecognize the modified version as still having the properties,characteristics and capabilities of the modified word or phrase. Ingeneral, but with the preceding discussion in mind, a numerical valueherein that is modified by a word of approximation may vary from thestated value by ±15%, unless expressly stated otherwise.

Embodiments of the present invention relate to conditioning polymericmedical devices, such as stents, after radiation sterilization, such aselectron beam (e-beam) sterilization, and after crimping the device ontoa delivery device, but before sterilization. The conditioning improvesthe product such that the product properties are in a narrower range.

Although the discussion that follows focuses on a stent as an example ofa medical device, the embodiments described herein are easily applicableto other medical devices, including implantable and insertable medicaldevices. More generally, embodiments of the present invention may alsobe applied to other devices, including, but not limited to,self-expandable stents, balloon-expandable stents, stent-grafts,vascular grafts, cerebrospinal fluid shunts, or generally tubularimplantable medical devices including catheters.

Stents are typically composed of a pattern or network of circumferentialand longitudinally extending interconnecting structural elements orstruts. The pattern or network of struts form the scaffolding, or thedevice body, of a stent. In general, the pattern of the struts isdesigned to contact the lumen walls of a vessel and to maintain vascularpatency. In general, a stent pattern is designed so that the stent canbe radially compressed (crimped) and radially expanded (to allowdeployment). Embodiments of the present invention are applicable tovirtually any stent design and are, therefore, not limited to anyparticular stent design or pattern. One embodiment of a stent patternmay include cylindrical rings composed of struts. The cylindrical ringsmay be connected by connecting struts.

FIG. 1 depicts an example of a stent 50 comprising a plurality ofinterconnected stent struts 52 configured to move relative to eachother. The stent struts 52 can be, for example, arranged in a sinusoidalor serpentine pattern. The stent struts 52 can form a plurality ofcircumferential rings 54 that may be arranged axially to form a tubularscaffold configured to support biological tissue after implantation ofthe stent. The rings may be connected by linking struts 56. There may beas few as one linking strut 56 per ring, but two, three, or more bepresent, or many more as depicted in FIG. 1. Although the cross-sectionof the struts in stent is shown as rectangular-shaped, the cross-sectionof the struts is not limited to those depicted, but may be circular,elliptical, or another cross-sectional shape.

All, substantially all, or some portion of the struts of the stentscaffolding can be made partially or completely from a biodegradablepolymer, a biostable polymer, or a combination thereof. In such a case,a scaffolding composed of a polymer, or primarily of a polymer, providesthe support or outward radial force to a vessel wall when implanted. Asused herein, the terms biodegradable, bioabsorbable, bioresorbable, andbioerodable are used interchangeably and refer to materials such aspolymers that are capable of being completely degraded and/or erodedwhen exposed to bodily fluids such as blood and can be graduallyresorbed, absorbed, and/or eliminated by the body. The processes ofbreaking down and absorption of the material can be caused by, forexample, hydrolysis and metabolic processes. Biostable refers tomaterials, such as polymers, that are not biodegradable.

As used herein, “polymeric stent” refers to a stent having a scaffoldingthat is made completely, or substantially completely, from a polymer, orthe scaffolding is made from a composition including a polymer andanother material. If the scaffolding is made from a compositionincluding a polymer and another material, the polymer is a continuousphase of the scaffolding, the scaffolding is at least 50% by weightpolymer, or the scaffolding is at least 50% by volume polymer. In someembodiments, a polymeric stent may have a scaffolding made from acomposition including a polymer and another material that is at least70%, at least 80%, at least 90%, or at least 95% by volume or by weightpolymer. Analogous definitions apply to a polymeric tube, or a polymericmedical device except that the reference to the scaffolding would bereplaced by “tube” for a polymer tube and “device body” for a medicaldevice. The “device body” of a medical device is the functional devicewithout a coating or layer of material different from that of which thedevice body is manufactured has been applied. If a device is amulti-layer structure, the device body is the layer(s) that form thefunctional device, and for a stent this would be the layer(s) whichsupport the bodily lumen.

The scaffolding of the embodiments of the present invention can be madein entirely from, or in part from one or a combination of biodegradablepolymers including, but not limited to, poly(L-lactide) (PLLA),polymandelide (PM), poly(DL-lactide) (PDLLA), polyglycolide (PGA), andpoly(L-lactide-co-glycolide). The tube or stent scaffolding can also bemade of any of the following: a random, alternating, or block copolymerof two or more of the above polymers; a random, alternating, or blockcopolymer of one or more of the above polymers, and one or more of thefollowing: polycaprolactone (PCL), poly(trimethylene carbonate) (PTMC),polydioxanone (PDO), poly(4-hydroxy butyrate) (PHB), poly ethyleneglycol (PEG), and poly(butylene succinate) (PBS); or any combinationthereof. The PLGA used can include any molar ratio of L-lactide (LLA) toglycolide (GA). In particular, the stent can be made from PLGA with amolar ratio of (LA:GA) including 85:15 (or a range of 82:18 to 88:12),95:5 (or a range of 93:7 to 97:3), or commercially available PLGAproducts identified as having these molar ratios. Combinations ofpolymers may be used and may be used in any proportions.

Polymers that may be used for struts of a bioabsorbable stent or othermedical device include semicrystalline biodegradable polymers, such asbiodegradable polyesters. In particular, the struts and/or scaffoldingcan be made substantially or completely out of biodegradable polyesters,or a composition including a biodegradable polyester, having a glasstransition temperature (Tg) above body temperature, which is about 37°C. for a human (but embodiments of the present invention also encompassother animals). For example, this includes PLLA and PLGA. In someembodiments, the polymer is one having a glass transition temperaturegreater than 25° C., 30° C., 35° C., 40° C., 45° C., or 50° C. In someembodiments, the Tg is not more than 70° C., 75° C., or 80° C.

Table 1 below provides the Tg for some of the above polymers.

TABLE 1 Glass transition temperatures of polymers. Glass-TransitionPolymer Temperature (° C.)¹ PGA 35-40 PLLA 60-65 PDLLA 55-60 85/15 PLGA50-55 75/25 PLGA 50-55 65/35 PLGA 45-50 50/50 PLGA 45-50 ¹MedicalPlastics and Biomaterials Magazine, March 1998.

An exemplary embodiment is a PLLA scaffolding with a coating includingPDLLA and everolimus. Another exemplary embodiment is a PLLA scaffoldingwith a coating including PDLLA and zotarolimus. Other exemplaryembodiments encompass a scaffolding of PLGA with a molar ratio of(LA:GA) 85:15 (or a range of 82:18 to 88:12), or of PLGA with a molarratio of (LA:GA) 95:5 (or a range of 93:7 to 97:3) with a coatingincluding PDLLA and everolimus and/or zotarolimus.

A few of the more important design characteristics of stents are radialor hoop strength, expansion ratio, coverage area, and longitudinalflexibility. Fracture toughness is also an important characteristic asit minimizes or reduces cracking as the stent is crimped for delivery,and/or expanded during deployment. Selection of materials and use ofspecific processing steps ensure that the design characteristics of thestent are met.

In some embodiments, a stent may be formed from a polymeric or metallictube by laser cutting the pattern of struts in the tube. Such tubes aretypically formed by methods such as, but not limited to, extrusion orinjection molding, as well as other conventional processes. The stentmay also be formed by laser cutting a metallic or polymeric sheet,rolling the pattern into the shape of the cylindrical stent, andproviding a longitudinal weld to form the stent. Other methods offorming stents are well known and include chemically etching a metallicor polymeric sheet and rolling and then welding it to form the stent.

In other embodiments, a metallic or polymeric filament or wire may alsobe coiled to form the stent. Filaments of polymer may be extruded, meltspun, solution spun or, eletrospun. These filaments can then be cut,formed into ring elements, welded closed, corrugated to form crowns, andthen the crowns welded together by heat or solvent to form the stent. Inother embodiments, the wires or filaments may be braided and/orinterwoven to form the scaffolding of the stent.

The manufacturing process may further include radially expanding thetube to an expanded diameter and cutting a stent pattern in the expandedtube. The tube is radially expanded to increase its radial strength,which can also increase the radial strength of the stent subsequentlyformed from the expanded tube. The radial expansion process tends topreferentially align the polymer chains along the radial or hoopdirection which results in enhanced radial strength. The radialexpansion step is crucial to making a stent scaffolding with thin strutsthat is sufficiently strong to support a lumen upon implantation. Inpreferred embodiments, the PLLA scaffolding is formed from an extrudedpolymer tube, which prior to expansion, the tube may be completelyamorphous or have a relatively low crystallinity, for example, less than20%, less than 10%, or less than 5%.

The tube may be radially expanded by heating the tube to a temperaturebetween Tg and the melting point of the polymer. At temperatures aboveTg, the amorphous domains of a substance, such as a polymer, change froma brittle vitreous state to a solid deformable or ductile state. Uponexpansion the tube is cooled to below the Tg of the polymer, typicallyto ambient temperature, where the tube is essentially maintained at anexpanded diameter due to the fact that below Tg the polymer changes to abrittle vitreous state. In other words, the Tg corresponds to thetemperature where the onset of segmental motion in the chains of thepolymer occurs. For example, for a PLLA tube, an expansion temperatureof 65-120° C. is preferred to optimize crystallinity and crystal size.

In addition to radial expansion, the tube can also be axially elongatedor extended, before, during, and/or after, but preferably during, theradial expansion process. The percent radial expansion is defined as RE%=(RE ratio−1)×100%, where the RE Ratio=(Inside Diameter of ExpandedTube)/(Original Inside Diameter of the tube). The percent of axialextension that the polymer tube undergoes is defined as AE %=(AEratio−1)×100%, where the AE Ratio=(Length of Extended Tube)/(OriginalLength of the Tube).

In preferred embodiments, the percent radial expansion may be betweenabout 200% and 500%, preferably 400% to 500%, or any specific valuewithin either of these ranges. The percent axial extension expansion maybe between about 20% and about 200%, preferably about 20% and about120%, or any specific value within either of these ranges. In preferredembodiments, the tube is a PLLA tube.

The tube may be optionally annealed after the radial and/or axialexpansion. The annealing may be performed to relief residual stresses,and/or to stabilize the polymer. Increasing the temperature of the stentor tube above ambient may cause radial shrinkage (a decrease indiameter) or, in general, changes in shape such as warping along theaxis of the stent or tube. This change in shape may be due to a releaseof residual stress that occurs during treatment. Thus, in furtherembodiments, radial shrinkage (strain recovery) or changes in shape canbe reduced or prevented by restraining the stent or tube duringtreatment by, for example and without limitation, placing the tube orstent over a mandrel.

Polymers are viscoelastic materials that exhibit stress relaxation andstrain recovery. When an ideal elastic material is subjected to a stresswhich is then released after a time period, the material recovers theoriginal shape. A rubber band is a close approximation of this idealperfectly elastic material. When an ideal viscous material is subjectedto a stress which is then released after a time period, the materialdoes not recover the original shape, but remains in the deformed shape.A liquid is an approximation of a purely viscous material. When aviscoelastic material is subjected to a stress which is released after atime period, the strain, that is the deformation, slowly recovers, butnot completely to its' initial shape. The slow recovery results in achange in dimensions over time. Stress relaxation occurs when aviscoelastic material is held at a constant strain, or constantdeformation, and the stress decreases, or “relaxes” over time as thepolymer chains rearrange.

Annealing also may impact the fracture toughness which is enhanced for asemicrystalline polymer by minimizing the size of crystalline domainsand achieving an optimal amorphous/crystalline ratio. The crystallinityprovides strength and stiffness (high modulus) to the polymer which isneeded for supporting a vessel. However, if the degree of crystallinityis too high, the polymer may be too brittle and is more susceptible tofracture. In some embodiments, the degree of crystallinity for a PLLAscaffolding may be about 10% to about 40%, or more narrowly, 30%-40%.

Annealing may also help stabilize the polymer and/or ameliorate theeffects of physical aging. Amorphous and semicrystalline polymersgenerally undergo physical aging during storage when the glasstransition of the amorphous region is greater than the storagetemperature. As shown in FIG. 2, a material that is amorphous orpartially amorphous and is above it glass transition temperature willexhibit an inflection point or departure from a linear relationshipbetween the specific volume and temperature as the polymer cools ifcooling is at a non-equilibrium rate. At some point during cooling, thepolymer chains do not have sufficient “time” to rearrange to theequilibrium state, and are thus trapped in a non-equilibrium state.After the inflection point which is the experimentally observed glasstransition temperature marked as Tg in FIG. 2, the specific volumefollows a different linear relationship with temperature. A polymer held(“aged”) at a temperature below the glass transition temperature willdensify over time if the rate of cooling was not an equilibrium rate.Cooling rates are almost always non-equilibrium rates in actualpractice. This densification is illustrated by the arrow shown in FIG.2. As the material densifies, the mechanical properties are alsochanging, generally becoming more brittle and less elastic. It is thisgradual change in mechanical properties and densification which isreferred to as the phenomenon of “physical aging.”

After cutting a stent pattern into the expanded tube, the stentscaffolding may then be optionally coated with a coating which caninclude a polymer and another substance such as a radiopaque agentand/or a drug. In some embodiments, the coating is polymer free.Typically, a coating on a stent may be formed by applying or depositinga coating composition including the polymer and/or drug, and/or othermaterials, dissolved, and/or dispersed, in a solvent on the surface ofthe stent scaffolding. The coating composition can be applied to a stentscaffolding by various methods, such as, dip coating, brushing, orspraying. Spray coating a stent typically involves mounting or disposinga stent on a support, followed by spraying a coating composition from anozzle onto the mounted stent in one or more passes with substantialsolvent removal between passes which may involve application of heat.

In order to make the stent ready for delivery, the stent is secured to adelivery balloon. In this process, the stent is compressed to a reduceddiameter or crimped over the balloon resulting in high stress and strainon the stent. Heating a stent during crimping can reduce or eliminateradially outward recoiling of a crimped stent which can result in anunacceptable profile for delivery. Crimping may also occur at an ambienttemperature. Therefore, crimping may occur at a temperature ranging from25° C. to 60° C., or higher, for a duration ranging from about 60seconds to about 5 minutes.

Once the stent has been crimped onto a support element, such as andwithout limitation, a catheter balloon, the stent delivery device ispackaged in sealed storage containers. Such containers are adapted toprotect the assembly from damage and environmental exposure (humidity,oxygen, light, etc.) which can have an adverse effect on the stent.Storage containers for a stent and delivery system can be designed to beany convenient form or shape that permits the effective enclosure of astent and delivery system assembly contained therein. A containerintended primarily to protect the stent and delivery system fromenvironmental exposure can be a pouch or sleeve. In preferredembodiments, the container is an aluminum based pouch and the pouch isfilled with Argon or another inert gas. Thus, the crimped stent ispackaged in an oxygen free environment.

After the stent is packaged, it is sterilized. Sterilization istypically performed on medical devices, such as stents and deliverysystems, to reduce the bioburden. Bioburden refers generally to thenumber of microorganisms with which an object is contaminated. Thedegree of sterilization is typically measured by a sterility assurancelevel (SAL) which refers to the probability of a viable microorganismbeing present on a product unit after sterilization. The required SALfor a product is dependent on the intended use of the product. Forexample, a product to be used in the body's fluid path is considered aClass III device.

Radiation sterilization is well known to those of ordinary skill theart. Medical articles composed in whole or in part of polymers can besterilized by various types of radiation, including, but not limited to,electron beam (e-beam), gamma ray, ultraviolet, infra-red, ion beam,x-ray, and laser sterilization. A sterilization dose can be determinedby selecting a dose that provides a required SAL. A sample can beexposed to the required dose in one or multiple passes.

The stent may be sterilized by exposure to electron beam (e-beam)radiation, or some other type of radiation. The radiation exposure canbe performed with a conventional e-beam radiation source. In someembodiments, the packaged stent may be exposed to a dose between 10-40,20-35, or 20-30 kGy. In other embodiments, the stent may be exposed to adose between 20-31 kGy or, more narrowly, 20-27.5 kGy.

Radiation exposure can degrade the properties of the polymers and drugs.In particular, the radiation can generate active species and inducechemical reactions in the polymer and drug. High-energy radiation suchas e-beam and gamma radiation tends to produce ionization and excitationin polymer molecules. These energy-rich species undergo dissociation,subtraction, and addition reactions that degrade the properties of apolymer in a sequence leading to chemical stability. The stabilizationprocess can occur during, immediately after, or even days, weeks, ormonths after irradiation which often results in physical and chemicalcross-linking or chain scission. Chain scission can result in areduction in molecular weight which can adversely affect mechanicalproperties, and, in the case of a degradable polymer, degradationproperties. In contrast, cross-linking would tend to increase themolecular weight of the polymer. Resultant physical changes can includeembrittlement, discoloration, odor generation, stiffening, andsoftening, among others.

Exposing a polymer to e-beam radiation causes the generation of freeradicals in the polymer. The degradation of polymer properties has beenassociated with free radical generation caused by the radiationexposure. Free radicals generated can become trapped within the polymer.The change in the polymer properties may continue as the trapped freeradicals continue to decay after the initial radiation exposure. “Freeradicals” refer to atomic or molecular species with unpaired electronson an otherwise open shell configuration. Free radicals can be formed byoxidation reactions. These unpaired electrons are usually highlyreactive, so radicals are likely to take part in chemical reactions,including chain reactions. The free radicals formed due to radiationexposure can potentially react with the polymer chains to cause chainscission. These reactions are dependent on e-beam dose (more generally,radiation dose), dose rate, as well as the environment of irradiation(such as e-beam environment) including the type of gas present, thehumidity, and the temperature.

The characteristics of the stents produced from the above process, thatis extrusion of a polymer tube which is axially and radially expandedand optionally annealed, laser cut to form a stent pattern, coated witha drug delivery coating including a polymer and a drug, crimped onto adelivery device, packaged in an inert gas, such as argon, and sterilizedby e-beam sterilization showed a drop in number-average molecular weightof the polymer of the scaffolding after sterilization. Aftersterilization, the molecular weight of the PLLA of a PLLA scaffolding islower than the molecular weight of the PLLA of the PLLA scaffoldingprior to sterilization. Thus, the sterilization process reduces themolecular weight of the PLLA polymer.

In addition, the presence of free radicals in the sterilized PLLA hasbeen monitored and the concentration of the free radicals is seen todecrease with time after e-beam exposure. The decrease in concentrationis believed to be primarily due to the termination of the free radicalsthrough reactions with the polymer chains which may result in chainscission. The concentration of free radicals does not decay to zerountil about 2 months under inert gas packaging condition.

Moreover, the PLLA stents produced as discussed above exhibited largevariations in the product properties from lot to lot. In particular, thefinal products exhibited a wide variation in the number-averagemolecular weight of the PLLA polymer of the polymeric scaffolding andthe radial strength of the polymeric stent as tested.

It was unexpectedly discovered that the polymeric stent continued tochange after the final operation of e-beam sterilization. Thus, theproperties of the polymeric stent scaffolding did not vary from oneprocessing run to another, but it was the experimental testing protocolthat varied. Specifically, due to the change in the properties withtime, a variation in the time from manufacture to evaluation resulted inproducts apparently exhibiting a large variation in product properties.With respect to the number average molecular weight of the PLLA polymer,after the initial decrease in the molecular weight of a PLLA stent thathas been sterilized with e-beam radiation, it was unexpectedly foundthat the molecular weight subsequently increases for a time period aftersterilization. FIG. 5 provides an example of the number averagemolecular weight of a PLLA polymer in a PLLA stent and a PLLA tube as afunction of time after sterilization by e-beam. The number averagemolecular weight of the polymer immediately after the e-beamsterilization is lower than that of the polymer prior to e-beamsterilization.

Unexpectedly, in addition to changes in the number average molecularweight, it was observed that the radial strength decreases with timeafter sterilization. FIG. 6 illustrates the radial strength with timeafter sterilization by e-beam.

Therefore, methods are needed to improve the product by providing amethod which results in a narrower range in the product properties, andparticularly a narrower range at product release. The variousembodiments of the present invention include methods of “conditioning” apackaged and sterilized stent, and or a crimped stent to obtain suchimprovement.

“Conditioning” a product may be exposing the product to an environmenthaving a controlled temperature of about 28° C. or a specific controlledtemperature above 28° C. for at least 0.5 hours. The exposure results inthe product, and as a result the materials of which the product isformed, being heated to the exposure temperature, and subsequently beingmaintained at the exposure temperature for a duration of time. Theduration of time of the exposure is greater than the duration of timeduring which the product is maintained at the specified controlledtemperature because it takes some time to heat the product to theexposure temperature. The exposure may occur after sterilization and/orafter crimping. “Conditioning” may be heating a product to a temperatureat or above 28° C. in a controlled environment, and subsequentlymaintaining the product at that temperature in the controlledenvironment for at least 0.5 hours. The controlled environment may beone in which the temperature is controlled to within a specified range,or one in which temperature, pressure and/or some other variable arecontrolled within a specified range. “Conditioning” may be exposing aproduct to an environment of controlled temperature at or above 28° C.,and potentially the environment is also at a controlled pressure, and/orother characteristics are controlled in the environment, such as %oxygen and/or humidity, for a duration of time sufficient to reach aspecified value of a property. The specified value may be a “stabilized”value of the property, that is a pseudo steady-state value, or it may bea partially stabilized value. In some embodiments, the specified valueis determined with reference to the pseudo-steady state or plateau valueof the property. As a non-limiting example, as shown in FIG. 6, theradial strength changes from the initial time to about 60 days where itreaches a pseudo-steady state or plateau value. The specified value maybe some specific percentage of the total change from the initial valueto the plateau value. For example, the specified value may be a valuewhich represents at least 25%, at least 30%, or at least 50% of thechange from the initial to the plateau value. As an exemplary andnot-limiting example, if the initial radial strength is 850 mmHg and theplateau value is 600 mmHG, then the total decrease is 250 mmHG, andtherefore, a value of 800 mmHg radial strength would represent a radialstrength decrease from the initial that is 20% of the total decrease(850−0.2*250)). For a polymeric stent, the initial value may be thevalue right before crimping, after crimping but before sterilization, orafter sterilization. The specified value may also be a percent increaseor decrease from the initial value without reference to the plateauvalue.

Without being bound by theory, it is believed that the exposure toincreased temperature chemically stabilizes the polymer of the stent andaccelerates the loss of free radicals. As discussed and shown below,exposing a PLLA stent to a temperature above ambient dramaticallyaccelerates the reduction in concentration of the free radicals afterradiation exposure. Also, without being bound by theory, it is believedthat the increased stent temperature resulting from the exposureincreases the rate of stress relaxation. The choice of an intermediatetemperature and time regimen for the conditioning results in an optimumvalue of partial stabilization of both properties, and helps to improvethe product because the properties of the polymeric stent fall within anarrower range after conditioning.

As discussed above, below Tg, polymer chains have very low mobility.Without being limited by theory, it is believed that when free radicalsthat are generated in a polymer that is well below its Tg, the freeradicals are trapped by polymer chains that have very low mobility, forexample, those chains near or at the amorphous-crystalline interface.However, it is believed that free radicals can be trapped even incompletely amorphous polymers with no crystallinity. The trapping offree radicals is typical for a polymer such as PLLA with a Tg above bodytemperature that is sterilized at or near ambient temperature. Since thefree radicals generated have very low mobility, the probability of freeradicals combining and terminating is relatively low due to their lowmobility. As the temperature of the polymer increases closer to or aboveTg, polymer chain mobility increases. The mobility of free radicalsincreases which increases the probability of self-terminating reactions.

Without being bound by theory, it is believed that the change in theradial strength observed after sterilization by radiation results fromstress relaxation. When the polymeric stent is crimped onto a deliverydevice, it is subjected to stress. As the stent is maintained in thecrimped state, the polymer chains slowly rearrange over time, and thus,after some time, the polymer chains are not in the same configuration asimmediately before the crimping operation. Without being bound bytheory, there may be some partial loss of chain orientation andresultant decrease in radial strength due to the chain rearrangement. Asdiscussed above, the mobility of the polymer chains increases as thetemperature of the polymer approaches and surpasses the Tg. Thus, therate at which the stress relaxes is a function of the temperature.

On the other hand, exposing the polymer to a temperature above Tg andbelow the melting temperature (Tm) or heating the polymer to atemperature above Tg and below Tm may result in changes in thecrystallinity, crystal size, and alignment of polymer chains. Therefore,exposure or heating above Tg may also result in undesirable changes inmicrostructure. Such changes may include an increase in crystal size anddegree of crystallinity and loss of radial alignment. Similarly,exposing the polymer to a temperature “close” to Tg or heating thepolymer to a temperature “close” to Tg, particularly for a sustainedperiod of time, may allow changes in the polymer, and in particular,loss of the chain orientation as the chains re-arrange to a more random,and thus more entropically favorable, configuration. Thus, theconditioning, that is the exposure to increased temperature withsubsequent stent heating, should be performed at a temperature and for aduration that inhibits loss, or inhibits significant loss, of mechanicalproperties generated by the radial and/or axial expansion and in laterpressing steps.

In some embodiments, the stent is conditioned after radiationsterilization, such as but not limited to e-beam sterilization, by beingheated as a result of exposure to an environment at a specifiedtemperature above ambient temperature, such as at or above 28° C., andleft in the environment for a duration of time not less than 0.5 hour.For example, the stent can be exposed in a temperature controlled ovenin which the temperature can be precisely controlled at a specifiedtemperature or within a temperature range.

Thus, in some embodiments, the exposure temperature, and thus thetemperature to which the stent is heated and maintained, can be atemperature that about 28° C., or in the range of 28° C. to about 15° C.below the Tg of the polymeric scaffolding of the stent. In preferredembodiments, the specific temperature of exposure is not more than about20° C. below the Tg of the polymeric scaffolding of the stent. In someembodiments, the specific temperature of exposure is not more than aboutis about 25° C. below the Tg of the polymeric scaffolding of the stent.

As used herein, the Tg of the polymeric scaffolding of the stent refersto the Tg of the scaffolding of the polymeric stent as measured bystandard differential scanning calorimetry (modulated or unmodulated)with a temperature ramp of 5-40° C./min and if modulated, with atemperature modulation of 0.01 to 10° C. with a modulation period of 1to 100 seconds, utilizing nitrogen or argon at a flow rate of 20-200ml/min, or as determined as the mid-point in the change of the heatcapacity in a plot of heat capacity versus temperature as illustrated inFIG. 7. FIG. 7 the heat capacity obtained via DSC versus temperature inthe region of the glass transition. As shown in FIG. 7, the heatcapacity at low temperatures below the glass transition is linear and asthe temperature increases there is a large increase in the slope withanother linear region followed by a decrease in slope at highertemperature. The temperature range of change from where slope changes atlow temperature, through the intermediate region to the change in slopeat higher temperatures is the glass transition region. The glasstransition temperature can be determined by extrapolating a line fitfrom the linear region at low temperature (line X), extrapolating a linefit from the linear region at high temperature (line Z), and a linefitted along the steep linear region (line Z) extrapolated to higher andlower temperatures so that line Z intersects line X at A and line Y atB. The temperature at A is called Tg,onset and the temperature at B isTg,end. The midpoint of Tg,onset and Tg along line Z (point C) isTg,midpoint which is calculated from Tg,onset+½ (Tg,onset+Tg,end). Somesources defined the Tg as Tg,onset while other sources define Tg asTg,midpoint. Unless otherwise specified, Tg refers to Tg,midpoint in thepresent application. Heat capacity determinations are typically moreaccurate when modulated DSC is used. The Tg that is measured is that ofpolymeric scaffolding in the “as processed condition,” which includesany other additives in the scaffolding, but does not include any coatingon the scaffolding. In other words, the Tg of the polymeric scaffoldingmay not be equivalent to the Tg of the polymer used in the fabricationof the scaffolding in the “as received” condition.

The polymeric scaffolding may be made from a blend of polymers, or acopolymer, either a random or block copolymer. In some polymer blendsand copolymers, and particularly for block copolymers, two or more glasstransition temperatures (Tgs) are observed. In general, the abovetemperature limits apply to the Tg that is above body temperature (about37° C. for a human being, but the embodiments of the present inventionare not limited to human beings as patients) if only one is above bodytemperature. In some embodiments, two or more Tgs may be above bodytemperature, and in such embodiments the specific temperature may be notmore than about 15° C. below the lowest Tg. In general, the specifictemperature may be not more than about 15° C. below the Tg of thepolymer which contributes most significantly to the radial strength andtoughness of the stent, but if two or more polymers have approximatelyequal contributions, the specific temperature is determined withreference to the lowest Tg of these two polymers. One of skill in theart is able to determine the relevant Tg(s) of the polymeric scaffoldingbased upon the disclosure herein.

The conditioning methods disclosed herein differ from merely placing asterilized and packaged polymeric stent into storage for some timeperiod. The standard United States Pharmacopeia (USP) storage conditionsallow for a broad range in temperature, such as from 15° C. to 30° C.Storage of a packaged stent under conditions that are not preciselycontrolled may result in product properties that fall within a narrowrange. In other words, if the conditions are not controlled, the productproperties may stabilize or change by different amounts resulting in abroader distribution of product properties at release. To improve theproduct by having product properties in a narrower range at productrelease, the conditioning of the product may be under controlledconditions. Thus, embodiments of the present invention include storageor exposure under conditions that are precisely controlled. In someembodiments, the specific temperature may be controlled to be within ±5°C., preferably within ±3° C., more preferably, ±2° C., and even morepreferably ±1.5° C.

In some embodiments, the stent is conditioned at, that is exposed to, aspecified temperature or temperature range for a duration of time,followed by reducing the temperature of exposure, for example, back toambient temperature.

In some embodiments, the specified temperature for conditioning, that isthe temperature to which the stent is exposed and/or the temperature towhich the stent is heated and at which it is subsequently maintained,for an arbitrary polymer with a Tg above body temperature (up to 15° C.below the Tg of the polymer) can be, in degrees Celsius, in the range of25-30, 30-35, 35-40, 40-45, 45-50, 50-55, 55-60, 60-65, 65-70, 70-75,75-80, 80-85, 85-90, 90-95, 95-100, or greater than 100. The specifiedtemperature for conditioning can be in ranges of 1 or 2 degrees Celsiusincrements from 25° C. to a temperature that is 15° C. below the Tg ofthe polymeric scaffolding. The temperature range may be 32° C.-40° C.,33° C.-37° C., or 35° C.-40° C. The temperatures above and disclosedelsewhere herein can also apply to the actual temperature of the stent.

The specified temperature for conditioning for PLLA in degrees Celsius,can be, between 25-30, 30-35, 35-40, 40-45, or 45-50. The specifiedtemperature for conditioning PLLA in degrees Celsius can also be in therange of 28-30, 30-32, 32-34, 34-36, 36-38, or 38-40. The specifiedtemperature for conditioning can be any temperature, in degrees Celsius,between 25-40.

The specified temperature for conditioning for 85/15 and 75/25 PLGA canbe in a temperature range, in degrees Celsius, of 25-30, 30-35, or35-40. The specified temperature for conditioning for 85/15 and 75/25PLGA can also be in a temperature range, in degrees Celsius, of 30-32,32-34, 34-36, 36-38, 38-40, 40-42, 42-44, 44-46, 46-48, or 48-50° C. Thespecified temperature for conditioning can be any temperature between25-50° C.

Embodiments of the present invention also encompass combinations of theabove ranges that result in a contiguous range. As a non-limitingexample, the following temperature ranges in degrees Celsius, 32-34,34-36, 36-38, and 38-40, can be combined to obtain the temperature range32-40 also expressed in degrees Celsius.

The duration of the conditioning, that is it may be the duration of theexposure at the specified temperature or it may be the duration that thestent is maintained at the specified temperature after heating to thespecified temperature, in combination with any of the disclosedtemperature embodiments, can be 4 hours to 20 days, less than 4 hours,or greater than 20 days. The duration of exposure or the duration ofstent heating, in combination with any of the disclosed temperatureembodiments, can be 0.5 hours to 1 hour, 0.5 hours to 2 hours, 0.5 hoursto 24 hours, 0.5 hours to 32 hours, 4 hours to 8 hours, 8 hours to 16hours, 8 hours to 18 hours, 8 hours to 20 hours, 8 hours to 36 hours, 4hours to 32 hours, 8 hours to 48 hours, 16 hours to 32 hours, 16 hoursto 48 hours, 16 hours to 72 hours, 36 hours to 120 hours, 36 hours to 4days, 0.5 days to 1 day, 0.5 day to 2 days, 0.5 day to 3 days, be 0.5day to 4 days, 0.5 day to 5 days, 0.5 to 10 days, 0.5 to 12 days, 0.5 to15 days, 0.5 to 20 days, 1 day to 2 days, 1 day to 3 days, 1 day to 4days, 1 day to 5 days, 1 day to 6 days, 1 day to 8 days, 1 day to 10days, 1 day to 12 days, 1 day to 15 days, 1 day to 20 days, 1 day to 21days, 2 days to 3 days, 2 days to 4 days, 2 days to 5 days, 2 days to 6days, 2 days to 7 days, 2 days to 8 days, 2 days to 9 days, 2 days to 10days, 2 days to 12 days, 2 days to 15 days, 2 days to 20 days, 3 days to4 days, 3 days to 5 days, 3 days to 6 days, 3 days to 7 days, 3 days to8 days, 3 days to 9 days, 3 days to 10 days, 3 days to 12 days, 3 daysto 15 days, 3 days to 20 days, 4 days to 5 days, 4 days to 6 days, 4days to 8 days, 4 days to 10 days, 4 days to 12 days, 4 days to 15 days,4 days to 16 days, 4 days to 20 days, 4 days to 5 days, 5 days to 6days, 5 days to 7 days, 5 days to 8 days, 5 days to 10 days, 5 days to15 days, 5 days to 18 days, 5 days to 20 days, 6 days to 8 days, 6 daysto 10 days, 6 days to 12 days, 6 days to 15 days, 6 days to 18 days, 6days to 20 days, 7 days to 8 days, 7 days to 9 days, 7 days to 10 days,7 days to 12 days, 7 days to 14 days, 7 days to 18 days, 7 days to 20days, 8 days to 9 days, 8 days to 10 days, 8 days to 12 days, 8 days to14 days, 8 days to 15 days, 8 days to 16 days, 8 days to 18 days, 8 daysto 20 days, 9 days to 10 days, 9 days to 11 days, 9 days to 15 days, 9days to 16 days, 9 days to 18 days, 9 days to 20 days, 10 days to 11days, 10 days to 12 days, 10 days to 14 days, 10 days to 15 days, 10days to 20 days, 10 days to 11 days, 10 days to 12 days, 10 days to 14days, 10 days to 15 days, 10 days to 16 days, 10 days to 18 days, 10days to 20 days, 11 days to 12 days, 11 days to 15 days, 11 days to 20days, 12 days to 13 days, 12 days to 14 days, 12 days to 15 days, 12days to 16 days, 12 days to 18 days, 12 days to 20 days, 15 days to 16days, 15 days to 18 days, 15 days to 20 days, 16 days to 18 days, 16days to 20 days, or 18 days to 20 days. Embodiments of the presentinvention also encompass durations in which each of the above lower andupper limits is preceded by “about.” For example, about 18 days to about20 days.

In preferred embodiments, the polymeric scaffolding is completely, oressentially completely, made of PLLA, PLGA with a molar ratio of (LA:GA)85:15 (or a range of 82:18 to 88:12), or PLGA with a molar ratio of(LA:GA) 95:5 (or a range of 93:7 to 97:3), the specific temperature ofthe conditioning is 30° C. to 35° C., controlled to within at least ±3°C. or at least ±2° C., and the conditioning duration is from 1 day to 12days, preferably, from 3 days to 8 days, and more preferably, from 5days to 7 days.

In other preferred embodiments, the polymeric scaffolding is completely,or essentially completely, made of PLLA, PLGA with a molar ratio of(LA:GA) 85:15 (or a range of 82:18 to 88:12), or PLGA with a molar ratioof (LA:GA) 95:5 (or a range of 93:7 to 97:3), the specific temperatureof the conditioning is in the range of 28° C. to 32° C., such as withoutlimitation, 30° C., and the specific temperature is controlled to withinat least ±3° C. or at least ±2° C., with a duration of the conditioningis from 0.5 day to 20 days, preferably, from 1 day to 15 days, morepreferably from 2 days to 10 days, and even more preferably from 3 daysto 8 days.

In other preferred embodiments, the polymeric scaffolding is completely,or essentially completely, made of PLLA, PLGA with a molar ratio of(LA:GA) 85:15 (or a range of 82:18 to 88:12), or PLGA with a molar ratioof (LA:GA) 95:5 (or a range of 93:7 to 97:3), the specific temperatureof the of the conditioning is in the range is in the range of 32° C. to37° C., such as without limitation, 35° C., controlled to within atleast ±3° C. or at least ±2° C., and for a duration of the conditioningthat is from 0.5 day to 15 days, preferably, from 1 day to 10 days, morepreferably from 1 day to 8 days, and even more preferably from 1 day to6 days.

In other preferred embodiments, the polymeric scaffolding is completely,or essentially completely, made of PLLA, PLGA with a molar ratio of(LA:GA) 85:15 (or a range of 82:18 to 88:12), or PLGA with a molar ratioof (LA:GA) 95:5 (or a range of 93:7 to 97:3), the specific temperatureof the conditioning is in the range of 30° C. to 36° C., such as withoutlimitation, 33° C., controlled to within at least ±3° C. or at least ±2°C., and for a duration of the conditioning that is from 0.5 day to 15days, preferably, from 1 day to 10 days, more preferably from 1 day to 8days, and even preferably from 1 day to 6 days.

In further embodiments, the polymeric stent may be conditioned in two“conditioning” stages, one after crimping but prior to sterilization andone after the sterilization. The conditioning stage after crimping maybe before or after packaging. For the conditioning stage after crimping,the specific temperature of the conditioning may be from about 32° C. toabout 15° C. higher than the glass transition temperature of thepolymer, and the duration of the exposure and/or heating and maintenancemay be at least 0.5 hours.

In some embodiments of conditioning after crimping and beforesterilizing, the specified temperature for conditioning for an arbitrarypolymer with a Tg above body temperature (up to 15° C. above the Tg ofthe polymer) can be, in degrees Celsius, in the range of 33-35, 35-40,40-45, 45-50, 50-55, 55-60, 60-65, 65-70, 70-75, 75-80, 80-85, 85-90,90-95, 95-100, or greater than 100. The specified temperature forconditioning can be in ranges of 1 or 2 degrees Celsius increments from35° C. to a temperature that is 15° C. above the Tg of the polymericscaffolding. In a preferred embodiments, the temperature for theconditioning may be not more than 10° C. above the Tg of the polymericscaffolding.

The specified temperature for conditioning for PLLA in degrees Celsius,can be, between 33-35, 35-40, 40-45, 45-50, 50-55, 55-60, 60-65, 65-70,70-75, or 75-80. The specified temperature for conditioning for PLLA indegrees Celsius can also be in the range of 32-34, 34-36, 36-38, 38-40,39-40, 40-42, 42-44, 44-46, 46-48, 48-50, 50-52, 52-54, 54-56, 56-58,58-60, 60-62, 62-64, 64-66, 66-68, 68-70, 70-72, 72-74, 74-76, 76-78 or,78-80. The specified temperature for conditioning can be anytemperature, in degrees Celsius, between 25-80.

In preferred embodiments of the conditioning stage after crimping butbefore sterilizing, the specific temperature for conditioning is in therange of 50° C. to 65° C., and the duration is from 1 hour to 32 hours,and may also include a conditioning stage after sterilizing where thespecific temperature is in the range of 33° C. to 37° C., and theduration is from 16 hours to 80 hours, 20 hours to 76 hours, or 24 hoursto 72 hours.

In preferred embodiments for the two stage conditioning, the polymericscaffolding is completely, or essentially completely, made of PLLA, PLGAwith a molar ratio of (LA:GA) 85:15 (or a range of 82:18 to 88:12), orPLGA with a molar ratio of (LA:GA) 95:5 (or a range of 93:7 to 97:3),and for the initial conditioning stage after crimping and beforesterilizing the specific temperature of the conditioning is in the rangeof 50° C. to 65° C., such as without limitation, 55° C., controlled towithin at least ±3° C. or at least ±2° C., and for a duration from 0.5hour to 32 hours, preferably, from 1 hour to 28 hours, and morepreferably from 1 hour to 24 hours. For the second conditioning stageafter sterilizing, the specific temperature of the conditioning is inthe range of 33° C. to 37° C., such as without limitation, 35° C.,controlled to within at least ±3° C. or at least ±2° C., and for aduration from 0.5 day to 10 days, preferably, from 1 day to 8 days, morepreferably from 1 day to 6 days, and even more preferably from 36 hoursto 80 hours.

In any of the embodiments, the conditioning can be performed by cyclingthe exposure temperature, and, thus the actual temperature of the stent.The temperature cycling can be performed by increasing the exposuretemperature, decreasing the exposure temperature, and then repeating theincreasing and decreasing one or more times. In such embodiments, theexposure temperature may be increased to a peak temperature followed bya decrease to a minimum temperature. The peak temperature and minimumtemperature can be the same every cycle or can be vary from cycle tocycle. The temperature may be held constant for a period of time at thepeak and/or minimum temperature. The hold time periods may vary witheach cycle, or may be constant for some or all cycles. In someembodiments, the duration of the exposure and/or heating may bedetermined by only including the time periods at the peak temperature.In alternative embodiments, the duration of the exposure and/or heatingmay be determined by including all of the time during the cycle. Instill other embodiments, the duration of exposure may be determined asthat portion of the cycle exceeding any specified temperature betweenthe minimum and peak temperatures. In some embodiments, both the peakand the minimum temperature are within the range of equal to or aboutequal to 25° C. to 15° C. below the Tg of the polymeric scaffolding, andthe duration is determined by including all time within the cycle. Ifthe exposure is by temperature cycling, the temperature is controlled bewithin ±5° C., preferably within ±3° C., more preferably, ±2° C., andeven more preferably ±1.5° C. of the set-point or the intendedtemperature profile. In embodiments with two stages of conditioning,neither, either one, or both stages may include temperature cycling. Ifboth stages include temperature cycling, the pattern and number ofcycles may be the same in each stage or different.

The temperature cycling exposure to the stent can be performed, forexample, by disposing the stent in a temperature-controlled oven. Theoven can be programmed to expose the stent to a selected time versustemperature profile.

In any of the conditioning embodiments, the controlled environment mayinclude controlling other variables aside from the temperature. Thecontrolled environment may be one where the pressure is at oneatmosphere. The controlled environment may be essentially humidity free(less than 5% relative humidity (rh), less than 1% rh, or less than 1000ppm water, where ppm may be by mass or by volume), or it may be a lowhumidity environment such for example and without limitation, between 5%and 15% rh, between 5% and 20% rh, or between 10% and 30% rh. Thecontrolled environment may be one which is free of oxygen, oressentially free of oxygen (less than 1000 ppm oxygen, or less than 100ppm oxygen where ppm may be by volume or by mass).

In still other embodiments, the polymeric stent can be crimped onto adelivery device at a specified temperature from 43° C. to 53° C., suchas in any of the following temperature ranges in degrees Celsius orcombinations of these temperature ranges: 43-44, 44-45, 45-46, 46-47,47-48, 48-49, 49-50, 50-51, 51-52, or 52-53. Other embodiments of thepresent invention encompass crimping at a temperature in the range of44° C. to 52° C., 45° C. to 51° C., 46° C. to 50° C., and 47° C. to 49°C. In still other embodiments, crimping occurs at about 48° C. For anyof the above crimping embodiments, the stent may also be conditionedafter sterilizing according to any of the above embodiments forconditioning after sterilization.

In preferred embodiments, the crimping is performed in any of the abovetemperature ranges, and the polymeric stent is conditioned aftersterilization where the specific temperature of the exposure and/or thespecific temperature to which the stent is heated and subsequentlymaintained is in the range of 33° C. to 37° C., such as withoutlimitation, 35° C., controlled to within at least ±3° C. or at least ±2°C., and for a duration from 0.5 day to 10 days, preferably, from 1 dayto 8 days, more preferably from 1 day to 6 days, and even morepreferably from 36 hours to 80 hours.

In some embodiments, crimping is performed in any of the abovetemperature ranges, and two stages of conditioning as recited in any ofthe above embodiments are preformed. However, in preferred embodiments,the polymeric stent is only conditioned after sterilizing, or thepolymeric stent is conditioned after sterilizing is performed and eithercrimped in the specified temperature range or conditioned after crimpingbut before sterilizing is performed, but not both.

In some embodiments, as a result of the conditioning, whether one-stageor two-stage, with or without temperature cycling, and with or withoutcrimping in a specified temperature range, the radial strength of thestent may be at least 10% lower than the initial radial strength, wherethe initial radial strength is that measured within about 8-12 hoursafter the completion of the e-beam sterilization for one stageconditioning, or within about 8-12 hours after the completion of thecrimping operation for two stage conditioning. In other embodiments, theradial strength of the stent may be 10% to 20% lower than the initialradial strength, 20% to 30% lower than the initial radial strength, 30%to 40% lower than the initial radial strength, or 20% to 40% lower thanthe initial radial strength.

In some embodiments, the change in radial strength is determined as apercent of the total decrease from the initial radial strength to theplateau radial strength which is illustrated by the value after about 60days in FIG. 6. In some embodiments, as a result of conditioning, theradial strength is at a value that represents about 20% to about 30% ofthe total decrease from the initial value to the plateau, a value thatrepresents about 20% to about 40% of the total decrease from the initialvalue to the plateau, a value that represents about 30% to about 50% ofthe total decrease from the initial value to the plateau, or a valuethat represents about 40% to about 60% of the total decrease from theinitial value to the plateau.

In some embodiments, as a result of the conditioning, whether one-stageor two-stage, with or without temperature cycling, and with or withoutcrimping in a specified temperature range, the number-average molecularweight of the polymer of the polymeric scaffolding may be at least 10%greater than the initial number-average molecular weight, where theinitial number-average molecular weight is that measured within about8-12 hours after the completion of the e-beam sterilization for eitherone stage or two-stage conditioning. In other embodiments, thenumber-average molecular weight of the polymer of the polymericscaffolding may be 10% to 20% greater than the initial number-averagemolecular weight, 20% to 30% greater than the initial number-averagemolecular weight, 20% to 40% greater than the initial number-averagemolecular weight, or 30% to 50% greater than the initial number-averagemolecular weight. For those embodiments in which the polymericscaffolding includes a polymer blend, the number-average molecularweight referred to is that of the polymer that contributes mostsignificantly to the radial strength and toughness of the stent, but, iftwo or more polymers have approximately equal contributions, thenumber-average molecular weight may be an average of the two or more(keeping in mind the correct methods of averaging polymer molecularweights). One of skill in the art is able to determine the relevantnumber-average molecular weight of the polymer of the polymericscaffolding based upon the disclosure herein.

In some embodiments, the change in number-average molecular weight ismeasured as a percent of the total increase from the initial value tothe pseudo-steady state or plateau value which is illustrated by thevalue after about 21 days in FIG. 5. In some embodiments, as a result ofconditioning, the number-average molecular weight is at a value thatrepresents about 20% to about 30% of the total increase from the initialvalue to the plateau, a value that represents about 20% to about 40% ofthe total increase from the initial value to the plateau, a value thatrepresents about 30% to about 50% of the total increase from the initialvalue to the plateau, or a value that represents about 40% to about 60%of the total increase from the initial value to the plateau.

In still further embodiments, the conditioning, whether one-stage ortwo-stage, with or without temperature cycling, and optionally combinedwith crimping in a specified temperature range, may be performed at aspecified temperature and for a duration such that the radial strengthof the polymeric stent is reduced by not less than 10%, between 10% and15%, between 10% and 15%, between 15% and 20%, between 20% and 25%, ormore than 25%, such as 25% to 50%, compared to the initial radialstrength.

In still further embodiments, the conditioning, whether one-stage ortwo-stage, with or without temperature cycling, and optionally combinedwith crimping in a specified temperature range, may be performed at aspecified temperature and for a duration such that number-averagemolecular weight of the polymer of the polymeric scaffolding isincreased by at least 10%, from about 10% to about 20%, from about 20%to about 30%, from about 30% to about 40%, or from about 40% to about60% as compared to the initial number-average molecular weight.

In still further embodiments, the conditioning, whether one-stage ortwo-stage, with or without temperature cycling, and optionally combinedwith crimping in a specified temperature range, may be performed at aspecified temperature and for a duration such that the radial strengthof the polymeric stent is at a value that represents about 20% to about30% of the total decrease from the initial value to the plateau, a valuethat represents about 20% to about 40% of the total decrease from theinitial value to the plateau, a value that represents about 30% to about50% of the total decrease from the initial value to the plateau, or avalue that represents about 40% to about 60% of the total decrease fromthe initial value to the plateau.

In still further embodiments, the conditioning, whether one-stage ortwo-stage, with or without temperature cycling, and optionally combinedwith crimping in a specified temperature range, may be performed at aspecified temperature and for a duration such that number-averagemolecular weight of the polymer of the polymeric scaffolding is at avalue that represents about 20% to about 30% of the total increase fromthe initial value to the plateau, a value that represents about 20% toabout 40% of the total increase from the initial value to the plateau, avalue that represents about 30% to about 50% of the total increase fromthe initial value to the plateau, or a value that represents about 40%to about 60% of the total increase from the initial value to theplateau.

DEFINITIONS

Ambient temperature can correspond to any temperature between 20° C. and25° C.

All ranges disclosed include endpoints of the ranges.

As used herein, a “polymer” refers to a molecule comprised of, actuallyor conceptually, repeating “constitutional units.” The constitutionalunits derive from the reaction of monomers. As a non-limiting example,ethylene (CH₂═CH₂) is a monomer that can be polymerized to formpolyethylene, CH₃CH₂(CH₂CH₂)_(n)CH₂CH₃ (where n is an integer), whereinthe constitutional unit is —CH₂CH₂—, ethylene having lost the doublebond as the result of the polymerization reaction. Althoughpoly(ethylene) is formed by the polymerization of ethylene, it may beconceptually thought of being comprised of the —CH₂— repeating unit, andthus conceptually the polymer could be expressed by the formulaCH₃(CH₂)_(m)CH₃ where m is an integer, which would be equal to 2n+2 forthe equivalent number of ethylene units reacted to form the polymer. Apolymer may be derived from the polymerization of two or more differentmonomers and therefore may comprise two or more different constitutionalunits. Such polymers are referred to as “copolymers.” “Terpolymers” area subset of “copolymers” in which there are three differentconstitutional units. The constitutional units themselves can be theproduct of the reactions of other compounds. Those skilled in the art,given a particular polymer, will readily recognize the constitutionalunits of that polymer and will equally readily recognize the structureof the monomer from which the constitutional units derive. Polymers maybe straight or branched chain, star-like or dendritic, or one polymermay be attached (grafted) onto another. Polymers may have a randomdisposition of constitutional units along the chain, the constitutionalunits may be present as discrete blocks, or constitutional units may beso disposed as to form gradients of concentration along the polymerchain. Polymers may be cross-linked to form a network.

As used herein, a polymer has a chain length of 50 constitutional unitsor more, and those compounds with a chain length of fewer than 50constitutional units are referred to as “oligomers.” As used todifferentiate between oligomers and polymers herein, the constitutionalunit will be the smallest unique repeating unit. For example, forpoly(lactide) the constitutional unit would be

even though the polymer may be formed by the reaction of the cyclicaldimer, lactide,

Similarly, for poly(ethylene) the constitutional unit used to count the“number” of constitutional units would be —CH₂— units, even thoughconventionally the constitutional unit is stated to be —CH₂CH₂— becauseit is always derived from the reaction of ethylene.

The “glass transition temperature,” Tg, is the temperature at which theamorphous domains of a substance, typically a polymer, change from abrittle vitreous state to a solid deformable or ductile state atatmospheric pressure. In other words, the Tg corresponds to thetemperature where the onset of segmental motion in the chains of thepolymer occurs. Furthermore, the chemical structure of the polymerheavily influences the glass transition by affecting mobility.

The “melting temperature,” T_(m), of a polymer is the temperature atwhich an endothermal peak is observed in a DSC measurement, and where atleast some of the crystallites begin to become disordered. The measuredmelting temperature may occur over a temperature range as the size ofthe crystallites, as well as presence of impurities and/or plasticizers,impacts the measured melting temperature of a polymer.

As used herein, a reference to the crystallinity of a polymer refers tothe crystallinity as determined by standard DSC techniques.

“Stress” refers to force per unit area, as in the force acting through asmall area within a plane. Stress can be divided into components, normaland parallel to the plane, called normal stress and shear stress,respectively. True stress denotes the stress where force and area aremeasured at the same time. Conventional or engineering stress, asapplied to tension and compression tests, is force divided by theoriginal gauge length.

“Strength” refers to the maximum stress along an axis which a materialwill withstand prior to fracture. The ultimate strength is calculatedfrom the maximum load applied during the test divided by the originalcross-sectional area.

“Radial strength” of a stent is defined as the pressure at which a stentexperiences irrecoverable deformation. The loss of radial strength isfollowed by a gradual decline of mechanical integrity.

“Modulus” may be defined as the ratio of a component of stress or forceper unit area applied to a material divided by the strain along an axisof applied force that results from the applied force. The modulus is theinitial slope of a stress-strain curve, and therefore, determined by thelinear hookean region of the curve. For example, a material has atensile, a compressive, and a shear modulus.

“Strain” refers to the amount of elongation or compression that occursin a material at a given stress or load, or in other words, the amountof deformation.

“Elongation” may be defined as the increase in length in a materialwhich occurs when subjected to stress. It is typically expressed as apercentage of the original length.

“Toughness” is the amount of energy absorbed prior to fracture, orequivalently, the amount of work required to fracture a material. Onemeasure of toughness is the area under a stress-strain curve from zerostrain to the strain at fracture. The stress is proportional to thetensile force on the material and the strain is proportional to itslength. The area under the curve then is proportional to the integral ofthe force over the distance the polymer stretches before breaking. Thisintegral is the work (energy) required to break the sample. Thetoughness is a measure of the energy a sample can absorb before itbreaks. There is a difference between toughness and strength. A materialthat is strong, but not tough is said to be brittle. Brittle substancesare strong, but cannot deform very much before breaking.

As used herein, a “drug” refers to a substance that, when administeredin a therapeutically effective amount to a patient suffering from adisease or condition, has a therapeutic beneficial effect on the healthand well-being of the patient. A therapeutic beneficial effect on thehealth and well-being of a patient includes, but it not limited to: (1)curing the disease or condition; (2) slowing the progress of the diseaseor condition; (3) causing the disease or condition to retrogress; or,(4) alleviating one or more symptoms of the disease or condition.

As used herein, a “drug” also includes any substance that whenadministered to a patient, known or suspected of being particularlysusceptible to a disease, in a prophylactically effective amount, has aprophylactic beneficial effect on the health and well-being of thepatient. A prophylactic beneficial effect on the health and well-beingof a patient includes, but is not limited to: (1) preventing or delayingon-set of the disease or condition in the first place; (2) maintaining adisease or condition at a retrogressed level once such level has beenachieved by a therapeutically effective amount of a substance, which maybe the same as or different from the substance used in aprophylactically effective amount; or, (3) preventing or delayingrecurrence of the disease or condition after a course of treatment witha therapeutically effective amount of a substance, which may be the sameas or different from the substance used in a prophylactically effectiveamount, has concluded.

As used herein, “drug” also refers to pharmaceutically acceptable,pharmacologically active derivatives of those drugs specificallymentioned herein, including, but not limited to, salts, esters, amides,and the like.

EXAMPLES

The examples set forth below are for illustrative purposes only and arein no way meant to limit the invention. The following examples are givento aid in understanding the invention, but it is to be understood thatthe invention is not limited to the particular examples. The parametersand data are not to be construed to limit the scope of the embodimentsof the invention.

Example 1: Free Radical Concentrations after E-Beam Irradiation

The following example illustrates the effect on free radicalconcentration of exposing stent made from a polymer to a temperatureabove ambient temperature after sterilization with radiation. The stentsused in the study are a scaffolding made from PLLA.

FIG. 3 and Table 2 depict the relative free radical concentration of thestent with heat treatment and the stent with no heat treatment as afunction of time after e-beam sterilization. The relative free radicalconcentration is the free radical concentration normalized to theinitial concentration immediately after e-beam sterilization. The stentsin the study were sterilized by e-beam radiation with a dose of 31 kGy.The stents were packaged in a foil pouch (MARVELSEAL™360—Nylon/Aluminum/Low Density PolyEthylene (LDPE)) made by Oliver-Tolasof Grand Rapids, Mich. The packages were sealed with an argon atmosphereinside.

TABLE 2 Relative Free Radical Concentration of stents without and withheat treatment after e-beam sterilization. No Heat Treatment 55° C. HeatTreatment Free Radical Free Radical Days Hours Concentration Days HoursConcentration 0 0 1.00 0 0 1.00 0.04 1 0.81 0.08 2 0.23 0.08 2 0.66 0.215 0.15 0.13 3 0.61 0.42 10 0.11 0.17 4 0.60 0.92 22 0.07 0.25 6 0.542.00 48 0.04 0.33 8 0.48 1.00 24 0.40 2.00 48 0.35 3.00 72 0.29 5.00 1200.26 7.13 171 0.23 30.21 725 0.09

Each data point for both no heat treatment and heat treatment aftere-beam exposure was generated by an individual packaged stent sample.The data for heat treatment was generated from stents subjected to aheat treatment in an oven for 2, 5, 10, 22, and 48 hours at 55° C. Thefree radical concentration for the stent samples not subjected to a heattreatment and the stent samples subjected to heat treatment was measuredusing Electron Spin Resonance (ESR), also known as Electron ParamagneticResonance (EPR), in Abbott Vascular, Temecula, Calif.

As shown by FIG. 3, the free radical concentration decays much fasterwith heat treatment than without. The free radical concentration isstill at about 0.09 at 35 days with no heat treatment while the freeradical concentration is less than half of that, 0.04, after only about2 days.

FIG. 4 also depicts the relative free radical concentration of the stentwith heat treatment and the stent with no heat treatment. The data forthe no heat treatment is the same as that in FIG. 5. One data point inFIG. 4 for heat treatment was generated from a stent subjected to a heattreatment in a oven for 5 hours at 55° C., which is from FIG. 3. Theadditional data points for the curve with heat treatment are predictedby pseudo first order decay kinetics with free radical concentration ofsample after 5 hours at 55° C. heat treatment. A comparison of FIG. 3and FIG. 4 shows that the kinetic model predicts a greater relative freeradial concentration with time than the experimental data. Therefore,the model may be used to select a desirable heat treatment temperature.For example, the relative free radical concentration can be measured forvarious temperatures and the decay profile may then be calculated fromthe single data points. The decay profiles may be expected to provide anupper bound to the decay of the free radial concentration vs. time forthe various temperatures.

Example 2: Number-Average Molecular Weight after E-Beam Irradiation

Polymeric stents manufactured from poly(L-lactide) (PLLA). Polymer tubeswere extruded from Poly(L-lactide) (PLLA), RESOMER® L 210 S, supplied byBoehringer Ingelheim by extrusion of a polymer tube, biaxially expandingthe polymer tube such that the radial expansion was within about 400% toabout 500% and the axial expansion was within about 20% to about 120%,and laser cutting a stent pattern into the tube to form the stent.Subsequently, radiopaque markers were placed in the stent, and then adrug delivery coating including a polymer and a drug (PDLLA andeverolimus) was applied. The stent was crimped onto a balloon catheterat a temperature of about 40° C. to about 55° C. The stents werepackaged in a foil pouch (MARVELSEAL™ 360—Nylon/Aluminum/LDPE) made byOliver-Tolas of Grand Rapids, Mich. The packages were sealed with anargon atmosphere inside. The stents in the study were sterilized bye-beam radiation with a dose of 31 kGy.

The number-average molecular weight was determined as a function of timeafter sterilization. The number-average molecular weight prior tosterilization was about 180 to about 200 KDaltons. Immediately followingsterilization, the number-average molecular weight was measured to beabout 78 KDaltons. As shown in FIG. 5, the number-average molecularweight increased with the time after sterilization reaching a plateau ofabout 115 KDaltons after about 21 days when stored at room temperature(25° C.±3° C.).

Example 3: Radial Strength after E-Beam Irradiation

Polymeric stents were manufactured from poly(L-lactide) (PLLA), coated,crimped, and sterilized as described in Example 2. The radial strengthwas determined by as the pressure at which irrecoverable deformation wasobserved. The testing utilized a MSI RX550™ Radial Force Testerequipment at a rate of 0.5 mm/second, and testing was conducted at 37°C. The radial strength of the polymeric stents were measured as afunction of time after sterilization. The stents were maintained at roomtemperature (25° C.±3° C.) from the time of sterilization tomeasurement. As shown in FIG. 6, the radial strength decreases as afunction of time after sterilization, reaching a “pseudo-steady state,”“pseudo-equilibrium,” or plateau value at about 60 days.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the claims are to encompasswithin their scope all such changes and modifications as fall within thetrue sprit and scope of this invention. Moreover, although individualaspects or features may have been presented with respect to oneembodiment, a recitation of an aspect for one embodiment, or therecitation of an aspect in general, is intended to disclose its use inall embodiments in which that aspect or feature can be incorporatedwithout undue experimentation. Also, embodiments of the presentinvention specifically encompass embodiments resulting from treating anydependent claim which follows as alternatively written in a multipledependent form from all prior claims which possess all antecedentsreferenced in such dependent claim (e.g. each claim depending directlyfrom claim 1 should be alternatively taken as depending from anyprevious claims).

1. (canceled)
 2. A method for conditioning a polymeric stent, the methodcomprising: selecting a specified value of a radial strength of apolymeric stent including a polymeric scaffolding to result fromexposing the polymeric stent to a controlled temperature for a durationof time, wherein the specified value is a 10% to 40% decrease from aninitial value of the radial strength of the polymeric stent; exposingthe polymeric stent to the controlled temperature greater than 28 deg C.for the duration of time of at least 30 minutes sufficient to reduce theradial strength of the polymeric stent to the specified value; whereinthe polymeric scaffolding is made of a polymer derived from reaction ofmonomers including L-lactide, and wherein the polymeric stent has beencrimped onto a delivery device, packaged, and sterilized prior to theexposure.
 3. The method of claim 2, wherein the polymeric scaffolding isformed from a polymeric tube that has been deformed by the applicationof stress, the deformation comprising radial expansion of the polymerictube at a temperature greater than that of the glass transitiontemperature of the polymeric tube.
 4. The method of claim 2, wherein thepolymeric scaffolding comprises a polymer selected from the groupconsisting of poly(L-lactide), poly(DL-lactide),poly(L-lactide-co-glycolide), and all combinations thereof in allproportions.
 5. The method of claim 2, wherein the scaffolding is madeof a random, alternating, or block copolymer of two or more of the groupof claim
 4. 6. The method of claim 2, wherein the exposure temperatureis not higher than 20° C. below the glass transition temperature of thepolymeric scaffolding.
 7. The method of claim 2, wherein the exposuretemperature is 30° C. to 40° C. and the duration of exposure is 1 day to20 days.
 8. The method of claim 2, wherein the specified value is a 10%to 20% decrease from the initial value of the radial strength.
 9. Themethod of claim 2, wherein the exposure temperature is controlled towithin ±3° C.
 10. The method of claim 2, wherein the specified value isa pseudo-steady state or plateau value.
 11. The method of claim 2,wherein the specified value is a 30% to 40% decrease from the initialvalue of the radial strength.
 12. The method of claim 2, furthercomprising determining the specified value of the radial strength of thepolymeric stent.
 13. A method for conditioning a polymeric stent, themethod comprising: selecting a specified value of a number averagemolecular weight (Mn) of a polymeric stent including a polymericscaffolding to result from exposing the polymeric stent to a controlledtemperature for a duration of time, wherein the specified value is a 10%to 60% increase from an initial value of the Mn of the polymeric stent;exposing the polymeric stent to the controlled temperature greater than28 deg C. for the duration of time of at least 30 minutes sufficient toincrease the Mn of the polymeric stent to the specified value; whereinthe polymeric scaffolding is made of a polymer derived from reaction ofmonomers including L-lactide, and wherein the polymeric stent has beencrimped onto a delivery device, packaged, and sterilized prior to theexposure.
 14. The method of claim 13, wherein the polymeric scaffoldingis formed from a polymeric tube that has been deformed by theapplication of stress, the deformation comprising radial expansion ofthe polymeric tube at a temperature greater than that of the glasstransition temperature of the polymeric tube.
 15. The method of claim13, wherein the polymeric scaffolding comprises a polymer selected fromthe group consisting of poly(L-lactide), poly(DL-lactide),poly(L-lactide-co-glycolide), and all combinations thereof in allproportions.
 16. The method of claim 13, wherein the exposuretemperature is not higher than 20° C. below the glass transitiontemperature of the polymeric scaffolding.
 17. The method of claim 13,wherein the duration of exposure is from about 2 days to about 6 days.18. The method of claim 13, wherein the exposure temperature is 30° C.to 40° C. and the duration of exposure is 1 day to 20 days.
 19. Themethod of claim 13, wherein the exposure temperature is controlled towithin ±3° C.
 20. The method of claim 13, wherein the specified value isa pseudo-steady state or plateau value.
 21. The method of claim 13,wherein the specified value is a 10% to 40% increase from the initialvalue of the Mn.
 22. The method of claim 13, further comprisingdetermining the specified value of the Mn of the polymeric stent.